Magnetic resonance imaging (MRI), also called nuclear magnetic resonance imaging (NMR imaging), is a non-invasive method for the analysis of materials and is used extensively in medical imaging. It is non-destructive and does not use ionizing radiation. In general terms, nuclear magnetic moments of nuclei in the imaged material are excited at specific spin precession frequencies, which are proportional to an external magnetic field. Radio frequency (RF) signals resulting from the precession of these spins are collected using receiver coils. By manipulating the magnetic fields, signals are collected that represent different regions of the volume under study. These signals are combined to produce a volumetric image of the nuclear spin density of the object.
In MRI, a body is subjected to a constant main magnetic field B0. Another magnetic field, in the form of electromagnetic radio frequency (RF) pulses, is applied orthogonally to the constant magnetic field. The RF pulses have a particular frequency that is chosen to affect particular nuclei (typically hydrogen) in the body. The RF pulses excite the nuclei, increasing the energy state of the nuclei. After the pulse, the nuclei relax and release RF energy as a free induction decay (FID) signal, which can be transformed into an echo signal. The echo signals are detected, measured, and processed into images for display. The RF pulses may have a broad frequency spectrum to excite nuclei over a large range of resonant frequencies, or the RF pulses may have a narrow frequency spectrum to excite a nuclei over a more narrow range of resonant frequencies.
Composite RF pulses may be used to excite nuclei over different ranges of resonant frequencies. In this manner, composite RF pulses may be transmitted to excite multiple ranges of resonant frequencies, thereby allowing for collection of received signals from multiple areas of interest, such as multiple slices, simultaneously.
When hydrogen nuclei relax, the frequency that they emit is positively correlated with the strength of the magnetic field surrounding them. For example, a magnetic field gradient along the z-axis, called the “slice select gradient,” is set up when the RF pulse is applied, and is shut off when the RF pulse is turned off. This gradient causes the hydrogen nuclei at the high end of the gradient (where the magnetic field is stronger) to precess at a high frequency (e.g., 26 MHz), and those at the low end (weaker field) to precess at a lower frequency (e.g., 24 MHz). When a narrow-banded RF pulse is applied, only those nuclei which precess at that particular frequency will be tilted, to later relax and emit a radio transmission. That is, the nuclei “resonate” to that particular frequency. For example, if the magnetic gradient caused hydrogen nuclei to precess at rates from 24 MHz at the low end of the gradient to 26 MHz at the high end, and the gradient were set up such that the high end was located at the patient's head and the low end at the patient's feet, then a 24 MHz RF pulse would excite the hydrogen nuclei in a slice near the feet, and a 26 MHz pulse would excite the hydrogen nuclei in a slice near the head. When a single “slice” along the z-axis is selected, only the protons in this slice are excited by the specific RF pulse to a higher energy level, to later relax to a lower energy level and emit a radio frequency signal.
The second dimension of the image is extracted with the help of a phase-encoding gradient. Immediately after the RF pulse ceases, all of the nuclei in the activated higher energy level slice are in phase. Left to their own devices, these vectors would relax. In MRI, however, the phase-encoding gradient (in the y-dimension) is briefly applied in order to cause the magnetic vectors of nuclei along different portions of the gradient to have a different phase advance. Typically, the sequence of pulses is repeated to collect all the data necessary to produce an image. As the sequence of pulses is repeated, the strength of the phase-encoding gradient is stepped, as the number of repetitions progresses. For example, the phase-encoding gradient may be evenly incremented after each repetition. These steps of the phase-encoding gradient are also referred to as “warp levels.” The number of repetitions of the pulse sequence is determined by the type of image desired and can be any integer, typically from 1 to 1024, although additional phase encoding steps are utilized in specialized imaging sequences. The polarity of the phase encoding gradient may also be reversed to collect additional RF signal data. For example, when the number of repetitions is 1024, for 512 of the repetitions, the phase encoding gradient will be positive. Correspondingly, for the other 512 repetitions, a negative polarity phase encoding gradient of the same magnitude is utilized.
After the RF pulse, slice select gradient, and phase-encoding gradient have been turned off, the MRI instrument sets up a third magnetic field gradient, along the x-axis, called the “frequency encoding gradient” or “read-out gradient.” This gradient causes the relaxing protons to differentially precess, so that the nuclei near one end of the gradient begin to precess at a faster rate, and those at the other end precess at an even faster rate. Thus, when these nuclei relax, the fastest ones (those which were at the high end of the gradient) will emit the highest frequency RF signals. The frequency encoding gradient is applied only when the RF signals are to be measured. The second and third dimensions of the image are extracted by means of Fourier transformation. Fourier transformation, or Fast Fourier transformation, permits the received RF signal to be decomposed into a sum of sine waves, each of different frequency, phases and amplitudes. For example,S(t)=a0+a1 sin(ω1+φ1)+a2 sin(ω2t+φ2)+
Alternatively, the amplitude of the received RF signal may be shown to decay exponentially, as represented by:
  A  =            A      0        ⁢          ⅇ                        -          t                          T          2                    
where t is time, A0 is the initial amplitude of the received signal, and the
  ⅇ            -      t              T      2      term is the decay constant that depends upon the uniformity of the main magnetic field, B0.
The Fourier transformation of the signal in the time domain can be represented in the equivalent frequency domain by a series of peaks of various amplitudes. In MRI, the signal is spatially encoded by changes of phase and frequency, which is then decomposed by performing a two-dimension Fourier transform to identify pixel intensities across the image.
While the z-axis was used as the slice-select axis in the above example, similarly, either the x-axis or y-axis may be set up as the slice-select axis depending upon the desired image orientation and the anatomical structure of the object of interest being scanned. For example, when a patient is positioned in the main magnetic field, one axis is utilized as the slice-select axis to acquire sagittal images, and another axis is utilized as the slice-select axis to acquire coronal images.
Regardless of the orientation of the selected scan, mathematically, the slice select gradient, phase-encoding gradient, and read-out gradient are orthogonal. The result of the MRI scan in the frequency domain representation (k-space) is converted to image display in the time domain data after a 2D or 3D Fast Fourier transform (FFT). Generally, in a transverse slice, the readout gradient is related to the kx axis and the phase-encoding gradient is related to the ky axis. In 3D MRI, an additional phase-encoding gradient is related to the kz axis to acquire data in a third dimension. When the number of phase-encoding levels is smaller than a binary number, the missing data may be filled with zeros to complete the k-space so that an FFT algorithm may be applied.
In k-space, data is arranged in an inhomogeneous distribution such that the data at the center of a k-space map contains low spatial frequency data, that is, the general spatial shape of an object being scanned. The data at the edges of the k-space map contains high spatial frequency data including the spatial edges and details of the object.
The more uniform the main magnetic field B0, and the more uniform the frequency of the gradients and RF pulses, the higher the resulting image quality, because the precessing nuclei become de-phased more quickly when subjected to non-uniform magnetic fields. The main magnetic field, the gradient magnetic fields, and the frequency composition of the RF excitation pulse may all cause quicker de-phasing if any of these elements are non-uniform.
In magnetic resonance imaging, for the same set of scan parameters, a shorter scan tends to reduce the signal-to-noise ratio (SNR), while a longer scan, which would have a correspondingly larger k-space map, tends to increase the signal-to-noise ratio as well as image quality. Ideally, a fast scan with a high signal-to-noise ratio is preferred.
The physical limitations, including signal-to-noise ratio (SNR) versus scan time, are balanced in a clinical environment, and MRI sequences are programmed to maximize image quality, including signal-to-noise ratios, image contrast, and the minimization of image artifacts. Attempts are made to minimize scan times, all the while minimizing the effects of any non-uniform magnetic fields in the main magnetic field, the gradient magnetic fields, and the RF pulse composition. Steady state free precession (SSFP) imaging sequences do not use a refocusing 180° RF pulse, and the data are sampled during a gradient echo, which is achieved by dephasing the spins with a negatively pulsed gradient before they are rephased by an opposite gradient with opposite polarity to generate the echo. Steady state free precession techniques often permit fast imaging with high signal-to-noise ratios, but they are susceptible to image artifacts due to inhomogeneities in the main B0 magnetic field.
In order to capitalize on the fast imaging times afforded by steady-state free precession imaging sequences, artifacts resulting from main B0 magnetic field inhomogeneities must be minimized. In this manner, fast scan times may be achieved with improved signal-to-noise ratios. However, none of the previous MRI imaging sequences and techniques provide adequate fast scan times, acceptable signal-to-noise measurements, and reduced-artifact images.
What is needed is a new type of MRI imaging sequence that provides acceptable fast scan times and signal-to-noise ratios, and eliminates steady-state free precession imaging artifacts.